The invention relates to a device for converting X-rays and gamma-rays to visible light. More particularly, this invention relates to an improved imaging device and method for converting low intensity X-rays and gamma-rays to visible light images capable of being viewed in a clinical or similar environment.
One of the great concerns in the use of X-ray and gamma-ray medical diagnosis is the biological damage produced by the high intensity flux required to achieve a good image in medical radiology. Even outside of medical applications; e.g., industrial and surveillance applications, high dosage requirements cause environmental and health problems. High dosage operational requirements are the limiting factor in the application of the computer assisted tomography (CAT) scanner and the use of fluoroscopes. In dental applications where X-ray exposures are most routine, sensitive organs, such as the thyroid and pituitary glands, are often accidentally and unnecessarily exposed to large doses of radiation from the presently utilized exta-oral X-ray machines which are not area selective. Current attempts toward alleviating high dosage requirements essentially consist of the following devices and approaches.
Various intensifying screens, films and combinations of these have been employed to reduce high dosage requirements. Rare-earth phosphor screens such as terbium activated gadolinium and lanthanum oxysulfide have high absorption efficiencies in the area of 60% at the typical medical X-ray energies of from 20 to 60 KeV. They also have high efficiencies in the subsequent conversion of the absorbed X-ray energy into large numbers of visible light photons. Proper coupling of the screens with films of high sensitivity in the band of the emitted visible light can be used to reduce the necessary exposure time by a factor of approximately 50 compared with the direct X-ray exposure of the film. However, for fluoroscopic examinations the screen alone is employed without film. In such cases, even dark-adapted eyes have difficulty distinguishing image details at normal X-ray doses. This factor, together with the required long exposure period, makes the radiation dosage unacceptably high.
Image intensifiers have also played a part in the quest to reduce required radiation dosages. X-ray image intensification began with a diode-type intensifier tube. In such a tube, kinetic energies in the order of tens of KeV are imparted to the photoelectrons generated either directly by X-rays or via X-visible-photocathode conversion before they impinge upon an output phosphor screen. Simultaneously, the electron image is also demagnified several times prior to arriving at the output screen. The demagnification in tandem with high photoelectron kinetic energies results in an intensified X-ray image. In these tubes, after the initial photoelectron generation, the number of electrons in the electron image remains constant, and is not multiplied. The electrical and electro-optical requirements of systems employing these tubes make the systems large, complex and cumbersome.
Recently, micro-channel plate (MCP) multipliers have been used directly as a photocathode for incident X-rays along with an output phosphor screen. The disadvantage of this approach is that the low probability of photoelectron production in MCP material at medical X-ray energies and the low probability of those photoelectrons that are produced deep in the material emerging and being multiplied, results in a quantum efficiency which is at most, a few percent. With such low efficiency there is loss of a great deal of information which can not be retrieved at later stages. Furthermore, X-rays which penetrate more than one channel before detection cause image degradation and loss of resolution.
In a second recent approach employing an MCP multiplier, visible light photocathode material is deposited directly on the back of an X-ray phosphor. The MCP multiplier follows the photocathode with its output phosphor. Incoming X-rays are first converted to visible light, the visible light is converted to photoelectrons, the electrons are amplified by the MCP multiplier and converted to visible light once again by the output phosphor. This second approach employing an MCP multiplier has a much higher quantum efficiency than the previously noted approach. The higher efficiency can be directly attributed to the employment of the X-ray phosphor which is highly efficient. However, the second approach exhibits an inherent problem. The proximity of the X-ray phosphor and the highly sensitive visible-light photocathode, required in the same vacuum envelope to preserve resolution, causes contamination of the photocathode material and severely limits the useful life of the intensifier. Most importantly, with both the foregoing approaches employing an MCP multiplier, X-rays must first enter the vacuum envelope of the intensifier before they are detected and intensified. Therefore, the X-rays must pass through a window material which seals the vacuum chamber. In order to prevent significant loss of quantum efficiency this window material must be very thin so that it is highly transparent to incident X-rays. On the other hand, it must also be thick enough to withstand a pressure differential of at least one atmosphere. These two basically conflicting requirements result in design compromises such as a curved window with a thickness in the order of a few hundred .mu.m which will, even at such minimal thicknessses, result in the loss of from 20% to 30% of the incident X-rays. Even so, while such thin windows may be adequate for the purposes of laboratory experimentation, they are generally too fragile for operation use.